Nuclear medicine diagnosis equipment

ABSTRACT

The present invention relates to a nuclear medicine diagnosis equipment comprising a scintillator block having a plurality of scintillators, the scintillator block having a plurality of scintillator arrays in a depth direction of an incident γ ray with different decay times for an emitted light pulse; an incidence timing calculating device for calculating an incident timing in the scintillator array; a scintillator array identifying device for identifying a scintillator array, in a plurality of arrays, that has received the electrical signal; and an incidence timing compensation device in a position arithmetic processing part for discriminating whether compensation for an incidence timing calculated by the incidence timing calculating device is to be done or not corresponding to a scintillator array identified by the scintillator array identification part.

FIELD OF THE INVENTION

The present invention relates to a nuclear medicine diagnosis equipment(ECT apparatus) for obtaining a tomogram image for a concerned region bydosing a subject with a radiopharmaceutical, and then by coincidencecounting of a pair of γ rays emitted from a positron emissionradioisotope (radioisotope, RI) accumulated in a region of interest ofthis subject, and the present invention specifically relates to atechnique for coincidence counting of the γ ray.

RELATED ART

A PET (Positron Emission Tomography) apparatus will be describedhereinafter as an example among the above-described nuclear medicinediagnosis equipments, i.e., ECT (Emission Computed Tomography)apparatuses. The PET apparatus has a configuration in which two γ raysthat are emitted from a region of interest of a subject in two mutuallyreverse directions making approximately 180 degrees are detected with γray detectors disposed facing each other, and a tomogram image of thesubject is reconstructed in simultaneous detection (coincidencecounting) of these γ rays. Furthermore, some of the γ ray detectors usedfor coincidence counting of a γ ray in the PET apparatus have ascintillator that emits light by incidence of a γ ray emitted from asubject, and a photo multiplier tube that converts the emitted lightinto an electrical signal in this scintillator.

Here, in principle, a γ ray emitted from a position distant from acenter of view often diagonally enters into the scintillator of the γray detector, and therefore the γ ray will be detected not only in thecorrect position, but also in an incorrect position. That is, parallaxerrors will gradually be larger toward a circumference part from acenter of view, resulting in inaccurate tomogram image obtained with thePET apparatus. For this reason, in some examples, the scintillator isdivided into a plurality of scintillators having different decay timesfor the emitted light pulse (in an optically combined condition) in anincident direction of the γ ray, for example, the scintillator isdivided into a scintillator array having a shorter decay time for the γray in an incidence side of the γ ray and a scintillator array having alonger decay time for the γ ray in a photo multiplier tube side. Therebythe position of the emitted γ ray will be detected with higher precisioneven in the case of diagonal incidence of the γ ray to the scintillatorof theγ ray detector, expecting improvement of acquisition of moreaccurate and precise tomogram images. (For example, refer to PatentDocuments 1, 2).

-   [Patent Document 1]-   JP-A, No. 06-337289 (pages 2 to 3, FIG. 1)-   [Patent Document 2]-   JP-A, No. 2000-56023 (pages 2 to 3, FIG. 1)

DESCRIPTION OF THE INVENTION Problems To Be Solved By The Invention

However, conventional nuclear medicine diagnosis equipments havefollowing problems. That is, scintillators having different decay timesfor an emitted light pulse often have different rise times for theemitted light pulse. Accordingly, a possible time lag will arise indetection between the opposing γ ray detectors in the case ofcoincidence counting with γ ray detectors using scintillator arrayshaving different rise times for an emitted light pulse. That is,actually simultaneously emitted γ rays may not be recognized, by thistime lag, to be simultaneously emitted γ rays, in coincidence countingprocessing based on detection with the γ ray detector, leading to apossible problem of reduction of detection sensitivity. Furthermore, inorder to improve reduction of the detection sensitivity, when acondition is set so that a judgment of being simultaneous may be givenin case of a detection between the opposing γ ray detectors having atime lag by setting wider a time range (timing window) for recognitionas an effective count in coincidence counting processing, influence ofrandom coincidence counting, scattered coincidence counting, etc. mayincrease to give degradation of reconstructed images, leading topossible problems.

The present invention has been made in view of the above-describedcircumstances, and aims at providing a nuclear medicine diagnosisequipment allowing acquisition of an accurate and precise tomogram imagehaving high sensitivity and providing avoidance of degradation of areconstructed image, even in use as a γ ray detector of a scintillatorhaving different decay times for an emitted light pulse.

Means For Solving The Problem

In order to achieve such objectives, the present invention has followingconfiguration. A nuclear medicine diagnosis equipment of the presentinvention comprises:

-   a scintillator block having a plurality of two-dimensionally and    closely arranged scintillators, the scintillator block having a    plurality of optically combined scintillator arrays in a depth    direction of an incident γ ray with different decay times for an    emitted light pulse;-   a photodetector for converting an emitted light pulse emitted in the    scintillator block into an electrical signal;-   an incidence timing calculating device for calculating an incident    timing in the scintillator array for the electrical signal outputted    from the photodetector;-   a scintillator array identifying device for identifying a    scintillator array, in a plurality of arrays, that has received the    electrical signal outputted from the photodetector; and-   an incidence timing compensation device for discriminating whether    compensation for an incidence timing calculated by the incidence    timing calculating device is to be done or not corresponding to a    scintillator array identified by the scintillator array    identification device, and for compensating the incidence timing    based on a result of discrimination.

Effectiveness of the present invention according to claim 1 will bedescribed hereinafter. First, a γ ray emitted from a subject enters intoa scintillator block having a two-dimensionally and closely arrangedplurality of scintillators and having a plurality of optically combinedscintillator arrays in a depth direction of an incident γ ray withdifferent decay times for an emitted light pulse. Furthermore, the γ raythat has entered into the scintillator block emits light in eachscintillator having different decay time for the emitted light pulse inthe scintillator array. Subsequently, the emitted light pulse that hasbeen emitted in each scintillator is converted into an electrical signalby a photodetector. In the next stage, an incidence timing calculatingdevice calculates a timing of incidence to the scintillator array forthe electrical signal outputted from the photodetector. Furthermore, ascintillator array identifying device identifies a scintillator arraythat has received the electrical signal outputted from thephotodetector, in a plurality of the scintillator arrays. An incidencetiming compensation device discriminates whether compensation for theincidence timing calculated by the incidence timing calculating deviceis to be done or not corresponding to the scintillator array identifiedby the scintillator array identification device, and compensates theincidence timing based on a result of discrimination. Accordingly, sincethe incidence timing compensation device discriminates whethercompensation for the incidence timing calculated by the incidence timingcalculating device is to be done or not corresponding to thescintillator array identified by the scintillator array identifyingdevice, and also compensates the incidence timing based on a result ofdiscrimination, a time lag of detection between the scintillator arrayscaused by difference of decay time of the emitted light pulse can becanceled by compensation even in case of coincidence counting by use ofa scintillator having different decay time to the emitted light pulse.Accordingly, a precise and accurate tomogram image accompanied byimprovement in detection sensitivity and by avoidance of degradation ofreconstructed image may be obtained.

Furthermore, the nuclear medicine diagnosis equipment according to claim2 of the present invention has an A/D converter for converting an analogsignal in a form of an electrical signal outputted from a photodetectorinto a digital signal, the scintillator array identifying devicecomprising:

-   an adding device for sequentially adding digital signals converted    by the A/D converter;-   an identified value calculating device for calculating an identified    value giving a value obtained by division of an intermediate added    value by a total added value, by using an intermediate added value    obtained, in the adding device, by addition of the digital signals    from a point of time of emission start of the emitted light pulse    that has been emitted in the scintillator block to a certain middle    point of time in the course of a point of time of emission end, and    a total added value obtained, in the adding device, by addition of a    digital signal from a point of time of emission start to a point of    time of emission end of the emitted light pulse that has been    emitted in the scintillator block in the adding device; and-   a discriminating device for discriminating whether the identified    value calculated by the identified value calculating device is a    larger value or a smaller value as compared with an intermediate    value between the identified values of each scintillator array    calculated by the identified value calculating device.

According to the nuclear medicine diagnosis equipment of claim 2 of thepresent invention, the A/D converter converts an analog signal in a formof an electrical signal outputted from the photodetector into a digitalsignal. Next, the adding device of the scintillator array identifyingdevice sequentially adds the digital signals converted by the A/Dconverter. Here, the identified value calculating device calculates theidentified value giving a value obtained by division of an intermediateadded value by a total added value, by using the intermediate addedvalue obtained, in the adding device, by addition of the digital signalsfrom a point of time of emission start to a certain intermediate pointof time that is an intermediate point of time from the point of time ofthe emission start to a point of time of the emission end of the emittedlight pulse in the scintillator block, and the total added valueobtained, in the adding device, by addition of a digital signal from apoint of time of emission start to a point of time of emission end ofthe emitted light pulse that has been emitted light in the scintillatorblock in the adding device. Furthermore, an intermediate valuecalculating device determines an intermediate value between theidentified values of each scintillator array calculated by theidentified value calculating device, and the discriminating devicediscriminates whether the identified value calculated by the identifiedvalue calculating device is a large value or a small value as comparedwith the intermediate value calculated by the intermediate valuecalculating device.

Accordingly, discrimination of whether the identified value calculatedby the identified value calculating device is a larger value or asmaller value may be done based on the sequential addition by the addingdevice of the scintillator array identifying device. That is, thescintillator array identifying device can identify which scintillatorarray in the scintillator has emitted the emitted light pulse.Furthermore, this equipment allows replacement of an integral actionconventionally performed by an integrator into an adding action of asequential addition in the adding device, leading to reduction of thenumber of parts, and cost reduction.

The nuclear medicine diagnosis equipment according to claim 3 of thepresent invention comprises:

-   a coincidence counting device for performing coincidence counting    using an incidence timing compensated by the incidence timing    compensation device and an incidence timing discriminated as not    compensated by the incidence timing compensation device;-   and-   a timing window storing device for storing a timing window showing a    predetermined range for performing coincidence counting by the    coincidence counting device as a timing window corresponding to a    combination of each of the plurality of scintillator arrays.

According to the nuclear medicine diagnosis equipment of claim 3 of thepresent invention, the coincidence counting device performs coincidencecounting using an incidence timing compensated by the incidence timingcompensation device and an incidence timing discriminated as notcompensated by the incidence timing compensation device. Here, thecoincidence counting is performed using a timing window showing apredetermined range that has been judged as the coincidence countingbeing coincident, and the coincidence counting is performed using atiming window corresponding to a combination of each of the plurality ofscintillator arrays stored by the timing window storing device.Accordingly, use of a different timing window corresponding to thecombination of each of the plurality of scintillator arrays allows ahigh-precision coincidence counting and reduction of influence of randomcoincidence counting and scattered coincidence counting, etc., leadingto high definition image with less noises.

Furthermore, the nuclear medicine diagnosis equipment according to claim4 of the present invention comprises:

a coincidence counting device for performing coincidence counting usingan incidence timing compensated by the incidence timing compensationdevice and an incidence timing discriminated as not compensated by theincidence timing compensation device; and

-   a timing window storing device for storing a timing window showing a    predetermined range for performing coincidence counting by the    coincidence counting device as a timing window corresponding to a    combination of each of the plurality of scintillators.

According to the nuclear medicine diagnosis equipment of claim 4 of thepresent invention, the coincidence counting device performs coincidencecounting using an incidence timing compensated by the incidence timingcompensation device and an incidence timing discriminated as notcompensated by the incidence timing compensation device. Here, thecoincidence counting is performed using a timing window showing apredetermined range that has been judged as the coincidence countingbeing coincident, and the coincidence counting is performed using atiming window corresponding to a combination of each of the plurality ofscintillators stored by the timing window storing device. Accordingly,use of a different timing window corresponding to the combination ofeach of the plurality of scintillators allows a high-precisioncoincidence counting and reduction of influence of random coincidencecounting and scattered coincidence counting, etc., leading to highdefinition image with less noises.

The nuclear medicine diagnosis equipment according to claim 5 of thepresent invention comprises a light guide for optically coupling thescintillator block and the photodetector.

According to the nuclear medicine diagnosis equipment of claim 5 of thepresent invention, the nuclear medicine diagnosis equipment comprisesthe light guide for optically coupling the scintillator block and thephotodetector. Accordingly, the light guide can suitably guide a lightfrom the scintillator block to the photodetector.

Furthermore in the nuclear medicine diagnosis equipment according toclaim 6 of the present invention, the a plurality of scintillator arraysis constituted by either one of scintillators of Gd₂SiO₅ (GSO) having Ceconcentration of 0.5 mol, Gd₂SiO₅ (GSO) having Ce concentration of 1.5mol, Lu₂SiO₅ (LSO), Lu_(x)Gd_(2-x)SiO₅ (LGSO), Lu_(x)Y_(2-x)SiO₅ (LYSO),Bi₄Ge₃O₁₂ (BGO), NaI, BaF₂, and CsF.

According to the nuclear medicine diagnosis equipment of claim 6 of thepresent invention, the a plurality of scintillator arrays is constitutedby Gd₂SiO₅ (GSO) having Ce concentration of 0.5 mol, Gd₂SiO₅ (GSO)having Ce concentration of 1.5 mol, Lu₂SiO₅ (LSO), Lu_(x)Gd_(2-x)SiO₅(LGSO) Lu_(x)Y_(2-x)SiO₅ (LYSO), Bi₄Ge₃O₁₂ (BGO), NaI, BaF₂, and CsF.Accordingly, various scintillators for constituting the plurality ofscintillator arrays may be selected, and thereby not only expensivescintillators but inexpensive scintillators may be used, leading toreduction of costs.

Furthermore, in the nuclear medicine diagnosis equipment according toclaim 7 of the present invention, the photodetector is made of a photomultiplier tube.

According to the nuclear medicine diagnosis equipment of claim 7 of thepresent invention, since the photodetector is a photo multiplier tube,it can appropriately convert a light from the scintillator block into anelectrical signal.

Moreover, in the nuclear medicine diagnosis equipment according to claim8 of the present invention, the photodetector is made of a photo-diode.

According to the nuclear medicine diagnosis equipment of claim 8 of thepresent invention, since the photodetector is a photo-diode, it canappropriately convert a light from the scintillator block into anelectrical signal.

Moreover, in the nuclear medicine diagnosis equipment according to claim9 of the present invention, the photodetector is an avalanchephotodiode.

According to the nuclear medicine diagnosis equipment of claim 9 of thepresent invention, since the photodetector is an avalanche photodiode,it can appropriately convert a light from the scintillator block into anelectrical signal.

Effect of the Invention

According to the nuclear medicine diagnosis equipment according to thepresent invention, the incidence timing compensation devicediscriminates whether the incidence timing calculated by the incidencetiming calculating device is to be compensated or not corresponding tothe scintillator array identified by the scintillator array identifyingdevice, and compensates the incidence timing based on the result of thediscrimination. Therefore, the difference of the time of detectionbetween the scintillator arrays caused by the difference of decay timeof the emitted light pulse may be canceled by compensation even in caseof coincidence counting using the scintillators having different decaytime for the emitted light pulse. Accordingly, a precise and accuratetomogram image providing a higher detection sensitivity and avoidance ofdegradation of a reconstructed image may be obtained.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram illustrating an entire configuration of a PETapparatus;

FIG. 2 is a block diagram illustrating a configuration of FPGA;

FIG. 3 is a perspective view illustrating a configuration of γ raydetector;

FIG. 4 is a graph illustrating a emitted light pulse of eachscintillator array outputted from an amplification circuit;

In FIGS. 5, (a) and (b) are views illustrating a timing of a γ ray thatenters into each scintillator array;

FIG. 6 is a graph illustrating added values from a point of time ofemission start of an emitted light pulse to a point of time of emissionend;

FIG. 7 is a graph for description of a timing window;

FIG. 8 illustrates a graph giving a timing spectrum without compensationof difference of detection time based on a different decay time of ascintillator array;

FIG. 9 illustrates a graph giving a timing spectrum with compensation ofdifference of detection time based on a different decay time of ascintillator array; and

FIG. 10 is a flow chart illustrating a signal processing after beingoutputted from an amplification circuit.

DESCRIPTION OF NOTATIONS

-   3 a, 3 b: A/D converter-   4 a, 4 b: Incidence timing calculation part (incidence timing    calculating device)-   6: Coincidence counting processing part (Coincidence counting    device)-   15: Scintillator block-   16: Light guide-   17: Photo multiplier tube (photodetector)-   18 a, 18 b: Scintillator array-   19 a, 19 b: Scintillator-   24: Scintillator array Identifying part (Scintillator array    identifying device)-   25: Adding part (Adding device)-   26: Identified value calculating part (Identified value calculating    device)-   28: Discriminating part (Discriminating device)-   29: Incidence timing compensation part (Incidence timing    compensation device)-   32: Timing window storing part (Timing window storing device)-   Tw: Timing window

BEST MODE FOR CARRYING OUT OF THE INVENTION

There has been realized an objective for obtaining a precise andaccurate tomogram image providing a higher detection sensitivity andavoidance of degradation of a reconstructed image even in use of ascintillator having different decay time for an emitted light pulse in aγ ray detector.

Embodiment

A PET (Positron Emission Tomography) apparatus will be described indetail, with reference to drawings. FIG. 1 is a block diagramillustrating a whole configuration of the PET apparatus. FIG. 2 is ablock diagram illustrating a configuration of FPGA7. In this embodiment,a PET apparatus is adopted as an example for a nuclear medicinediagnosis equipment, and description will be given.

The configuration of the whole PET apparatus will be described withreference to FIG. 1. As illustrated in FIG. 1, the PET apparatuscomprises γ ray detectors 1 for outputting an electrical signal afterconversion of a light emitted by a positron emission radioisotope(radioisotope, RI) that has been given to a subject M in a form ofradiopharmaceuticals and has been accumulated in a region of interest ofthis subject M. Each of the γ ray detectors 1 is disposed without anyclearance in a circumference of a body axis of the subject M, forexample, in a shape of a ring having a size about 700 mm in diameter(only two γ ray detectors 1 are illustrated in FIG. 1). Accordingly, twoγ rays emitted in opposite directions making approximately 180 degreesfrom the region of interest of the subject are detected by the γ raydetectors 1 facing each other to be outputted after converted intoelectrical signals.

Furthermore, the PET apparatus comprises: amplification circuits 2 a and2 b for amplifying electrical signals outputted from the γ ray detectors1; A/D converters 3 a and 3 b for converting analog signals amplified inthese amplification circuits 2 a and 2 b into digital signals; incidencetiming calculation parts 4 a and 4 b for receiving the electricalsignals amplified in the amplification circuits 2 a and 2 b and forcalculating incidence timings of the γ rays detected with the γ raydetectors 1; a position arithmetic processing part 5 for calculating aposition of the γ ray detector 1 that has received the γ ray emittedfrom the subject M based on the digital signals converted by these A/Dconverters 3 a and 3 b; and a coincidence counting processing part 6 forperforming processing for detection of coincident incidence of the γ ray(coincidence counting) in these two γ ray detectors 1 based oninformation from the position arithmetic processing part 5 and theincidence timing calculation parts 4 a and 4 b. Here, as illustrated inFIG. 2, the position arithmetic processing part 5 and the coincidencecounting processing part 6 are included in one programmable LSI (largescale integration circuit) called an FPGA (Field Programmable GateArray)7. In addition, the FPGA7 has functions, such as CPU 8, ROM 9, andRAM 10, and the position arithmetic processing part 5 and thecoincidence counting processing part 6 are one of functions of the CPU 8in the FPGA 7. Furthermore, as illustrated in FIG. 1, the FPGA 7 has areconstruction part 11 for reconstructing a tomogram image of thesubject when it is discriminated that a γ ray has been coincidentlydetected (coincidence counting) in the coincidence counting processingpart 6.

In addition, the apparatus of this embodiment is provided with acontroller 12, a monitor 13, an input part 14, etc. Hereinafter, theconfiguration of each part of the apparatus of this embodiment willspecifically be described.

The configuration of the γ ray detector 1 will be described using FIG.3.

FIG. 3 is a perspective view illustrating a configuration of the γ raydetector 1. As illustrated in FIG. 3, the γ ray detector 1 has thescintillator 19 divided and disposed also in a depth direction of the γray incidence, that is, the γ ray detector 1 is a DOI (Depth OfInteraction) detector having scintillators disposed in three dimensions.For example, this DOI detector is configured with a scintillator block15, a light guide 16, and a photo multiplier tube (PMT) 17.

In the scintillator block 15, two scintillator array 18 a andscintillator array 18 b having different decay time for an emitted lightpulse in a depth direction (a direction Z) of an incident γ ray areoptically coupled. The scintillator array 18 a has a plurality ofscintillators 19 a, and the scintillator array 18 b has a scintillator19 b in a closely contact arrangement in two dimensions, respectively.In detail, in the scintillator block 15, the scintillator array 18 ausing the scintillator 19 a (for example, Lu_(x)Y_(2-x)SiO₅ (LYSO))having a shorter decay time for an emitted light pulse in an incidenceside of the γ ray (front step), and the scintillator array 18 b usingthe scintillator 19 b (for example, Gd₂SiO₅ (GSO) having Ceconcentration of 0.5 mol) having a longer decay time for the emittedlight pulse in a light guide 16 side (back step) are stacked in twosteps (two pieces). In these scintillators 19 a and 19 b, light isemitted in response to a γ ray emitted from the subject M. At thispoint, since the scintillators 19 a and 19 b have decay times differentfrom each other for the emitted light pulse, they have different risetimes. A longer decay time will give a later rise time, resulting indifference between the detection time of the scintillator 19 a and thatof the scintillator 19 b. Here, the two scintillator array 18 a and 18 bare formed, respectively, with the scintillators 19 a, 19 b in a shapeof 8×8 chips (direction of X, direction of Y). Furthermore, lightreflecting materials and optical transmitting materials, and opticaladhesives for proportional distribution of light generated by incidenceof a γ ray in a direction of X and a direction of Y are inserted orcharged between the scintillators 19 a adjacent each other within thescintillator arrays 18 a and 18 b and between the scintillators 19 baccording to the position thereof.

The light guide 16 guides a light generated in the scintillators 19 aand 19 b of the scintillator block 15 to the photo multiplier tube 17.The light guide 16 is inserted between the scintillator block 15 and thephoto multiplier tube 17, and is optically coupled with each other usingthe optical adhesive, respectively.

The photo multiplier tube 17, for example, has 4 (channel) photoelectricconversion plates built therein. A light generated in the scintillators19 a and 19 b enters into four PMT photoelectric conversion plates, iselectronically amplified, and then finally is converted into anelectrical signal (analog signal) to be outputted. Accordingly, anoutput of this photo multiplier tube 17 forms an output of the γ raydetector 1. The above-described photo multiplier tube 17 is equivalentto a photodetector.

An incidence timing calculation parts 4 a and 4 b will be described withreference to FIG. 1, FIG. 4 to FIG. 5( b). FIG. 4 is a graphillustrating an emitted light pulse of the scintillator arrays 18 a and18 b outputted from the amplification circuit 2 a or amplificationcircuit 2 b. FIG. 5( a) and FIG. 5( b) are figures illustrating timingsof a γ ray that entered into the scintillator arrays 18 a and 18 b.Here, curves of (A) in FIG. 4 and FIG. 5( a) illustrate curves of the γray that has entered into the scintillator array 18 a with a shorterdecay time for the emitted light pulse, and curves of (B) illustrate acurve of the γ ray that has entered into the scintillator array 18 bwith a longer decay time for the emitted light pulse.

As illustrated in FIG. 1, an electrical signal outputted from the γ raydetector 1 is inputted into the incidence timing calculation parts 4 aand 4 b through the amplification circuits 2 a and 2 b. The incidencetiming of the γ ray that has entered into the scintillator arrays 18 aand 18 b of the γ ray detector 1 is calculated based on this electricalsignal. In detail, the incidence timing calculation parts 4 a and 4 bhave what is called a wave height and rise time compensation circuit,that is, an ARC (Amplitude and Rise-time Compensation) 20, and a timinggenerator circuit 21.

Analog signals, with different decay times for the emitted light pulse,that have been outputted from the amplification circuits 2 a and 2 b,for example, as illustrated in FIG. 4, are inputted into the ARC20 basedon the γ ray that has entered into the scintillator arrays 18 a and 18 bof the γ ray detector 1. Furthermore, the ARC20 performs waveformshaping processing for each of these analog signals in order tocalculate the incidence timing of the γ rays that have entered into thescintillator arrays 18 a and 18 b. In detail, this waveform shapingprocessing may be performed by an addition operation of values obtainedby delay processing of signals obtained from the amplification circuits2 a and 2 b and values obtained by a reversal processing and reductionprocessing of values of voltage of signals obtained from theamplification circuits 2 a and 2 b. Thus, the waveform is shaped into avoltage waveform as illustrated in FIG. 5( a). Here, t_(SF), t_(SR) thatare points of time (zero cross point) giving a voltage of 0 will givethe incidence timings of the γ rays that have entered into thescintillator arrays 18 a and 18 b. Furthermore, the timing generatorcircuit 21 converts the signal illustrating the incidence timingcalculated by the ARC20, as illustrated in FIG. 5( b), into a digitalsignal, and the signal is temporarily stored in an incidence timingstorage part 22 that is one function of a RAM10 of the FPGA7. Here, theabove-described incidence timing calculation parts 4 a and 4 b areequivalent to the incidence timing calculating device.

The position arithmetic processing part 5 will be described withreference to FIG. 1. As illustrated in FIG. 1, electrical signalsoutputted from the γ ray detector 1 are inputted through theamplification circuits 2 a and 2 b and furthermore, digital signalsconverted by the A/D converters 3 a and 3 b from the analog signalsinputted from these amplification circuits 2 a and 2 b for constantlyperforming A/D conversion are temporarily stored in an A/D convertedsignal storage part 23 that is one function of the RAM10 of the FPGA7.The position arithmetic processing parts 5 a and 5 b perform operationprocessing for determining the positions of the scintillators 19 a and19 b of the γ ray detector 1 that have received the γ rays emitted fromthe subject M based on the digital signals stored in this A/D convertedsignal storage part 23. Here, the operation of the position is performedusing values of voltage based on distribution of light to four inputPMTs located in the back step of the scintillators 19 a and 19 b of theγ ray detector 1 in an X direction and a Y direction (within the samescintillators 18 a and 18 b) of the scintillators 19 a and 19 b of the γray detector 1 (Anger logic).

Furthermore, as illustrated in FIG. 2, the position arithmeticprocessing parts 5 a and 5 b have a scintillator array identificationpart 24 for identifying which scintillator arrays 18 a and 18 b of thetwo scintillators have detected the incident γ rays with the γ raydetector 1. In other words, identification of this scintillator array 18a or scintillator array 18 b will perform an operation of position forthe Z direction giving information of which of the scintillator 19 a orthe scintillator 19 b of the γ ray detector 1 has received the γ ray.

The scintillator array identifying part 24 will be described withreference to FIG. 6. FIG. 6 is a graph illustrating an added value froma point of time of emission start to a point of time of emission end T2of the emitted light pulse. A curve of (A) in FIG. 6 illustrates a curveof a γ ray that has entered into the scintillator 19 a (the scintillatorarray 18 a) with a shorter decay time for the emitted light pulse, and acurve of (B) illustrates a curve of the γ ray that has entered into thescintillator 19 b (the scintillator array 18 b) with a longer decay timefor the emitted light pulse. The scintillator array identifying part 24comprises:

-   an adding part 25 for adding sequentially digital signals converted    by A/D converters 3 a and 3 b; an identified value calculating part    26 for calculating an identified value giving (intermediate added    value A_(T1))/(total added value A_(T2)), (division of intermediate    added value AT₁ by total added value A_(T2)), using an intermediate    added value A_(T1) obtained, in the adding part 25, by addition of    the digital signals from a point of time of emission start to a    certain intermediate point of time T₁, T₁ being an intermediate    point of time from the point of time of the emission start to a    point of time of the emission end T₂ of the emitted light pulse in    the scintillator block 19 a and 19 b, and a total added value A_(T2)    obtained by addition of the digital signals from a point of time of    emission start to a point of time of emission end of the emitted    light pulse that has been emitted in the scintillator block 19 a and    19 b; and a discriminating part 28 for discriminating, based on    emitted light pulses that have been emitted in two scintillators 19    a and 19 b, whether the identified value calculated by identified    value calculating part 26 is a larger value or a smaller value as    compared to an intermediate value K between identified values from    each scintillator array calculated by the identified value    calculating part 26.

Accordingly, the scintillator array identifying part 24 can identifywhich scintillator array of the two scintillators 18 a and 18 b hasdetected the incident γ ray with the γ ray detector 1 based on adiscriminated result in the discriminating part 28. When the result ofcalculated A_(T1)/A_(T2) is larger than the intermediate value K, thescintillator 19 a with a shorter decay time will be identified, andconversely, when it is smaller, the scintillator 19 b with a longerdecay time will be identified. The intermediate value K is anintermediate value between both peak values (values of voltage) whereinA_(T1), that is a value at most distant point of time Fs×m in waveformsof two patterns, is set in the addition process in the adding part 25(Fs×m: Fs is a sampling interval of A/D conversion and m is a number oftimes of addition). The intermediate value K is beforehand acquired byexperiments as data for discrimination, and is stored as an intermediatevalue data table 27, which is one function of the ROM9 of the FPGA7. Theabove-described scintillator array identifying part 24 is equivalent tothe scintillator array identifying device. The above-described addingpart 25 is equivalent to the adding device. The above-described theidentified value calculating part 26 is equivalent to the identifiedvalue calculating device. The above-described discriminating part 28 isequivalent to the discriminating device, and an intermediate valuestored by the intermediate value data table 27 is read out at the timeof discrimination processing.

The position arithmetic processing parts 5 a and 5 b have an incidencetiming compensation part 29 for discriminating whether the incidencetimings calculated by the incidence timing calculation parts 4 a and 4 bare to be compensated, corresponding to the scintillator arrays 18 a and18 b identified by the scintillator array identifying part 24, and forsubsequently compensating the incidence timings based on the results ofdiscrimination. In detail, when one scintillator of the scintillatorarray identified by the scintillator array identifying part 24 isidentified as the scintillator 19 b with a longer decay time, anoperation processing for setting an incidence timing t_(SR) calculatedby the incidence timing calculation parts 4 a and 4 b as t_(SR)−Δt(incidence timing compensated value) is performed. Subsequently, theresult is outputted to post-compensation incidence timing storage part30 that is one function of the RAM10 of the FPGA7. Conversely, when onescintillator of the scintillator array identified by the scintillatorarray identifying part 24 is identified as the scintillator 19 a with ashorter decay time, no compensation to an incidence timing t_(SF) isgiven. Then, the original incidence timing t_(SF) without compensationinputted into the scintillator array identifying part 24 is outputted tothe post-compensation incidence timing storage part 30. Furthermore, thepost-compensation incidence timing storage part 30 temporarily storesthe incidence timing t_(SF) and the incidence timing t_(SR) aftercompensation of difference of detection time based on a differencebetween decay times of scintillator array 18 a and of scintillator array18 b, in a relationship between an incidence timing t_(SF) and anincidence timing t_(SR). The above-described incidence timingcompensation part 29 is equivalent to the incidence timing compensationdevice.

In addition, this Δt (incidence timing compensated value) is beforehandacquired by experiments as data for compensation, and a difference oftime of rise times between the scintillator array 18 a and thescintillator array 18 b is stored as an incidence timing compensatedvalue in a compensation data table 31 that is one function of the ROM9of the FPGA7. Here, an incidence timing compensated value stored in thecompensation data table 31 in compensation processing is read out in theincidence timing compensation part 29.

The coincidence counting processing part 6 reads incidence timingst_(SF) and t_(SR) that have been stored in the post-compensationincidence timing storage part 30 every predetermined period of time (forexample, 128 ns), and performs coincidence counting of incidence timingst_(SF) and t_(SR), in this case, 4 combinations, from the 2 γ raydetectors 1. When the difference of time calculated by these 4combinations is within a timing window Tw (for example, 6 ns) that is apredetermined time range, the coincidence counting processing part 6discriminates the coincidence count as a valid coincidence count, andwhen not, it discriminates the coincidence count as an invalid count.

Incidentally, the timing window Tw will be described using FIG. 7, incase of coincidence counting processing of a γ ray that enters into eachγ ray detector 1 having one layer (piece) of scintillator array. FIG. 7is a graph for describing the timing window Tw. An ordinate axis A givesa number of times of an event (γ ray detection by coincidence countingprocessing), and an abscissa axis T gives a difference of time ofdetection of a γ ray by coincidence counting processing. In this way, atiming spectrum as illustrated in FIG. 7 is obtained. In this timingspectrum, a case where the abscissa axis T is 0 (detection of γ ray doesnot have difference of time) gives many number of times of events, andthe larger difference of time, the smaller number of times of events.That is, when the abscissa axis T gives 0, this spectrum illustrates agraph similar to a Gaussian distribution that the number of times ofevents of the ordinate axis A gives a peak. Here, an intermediate value(A/2) of the ordinate axis A in the timing spectrum illustrated in FIG.7 is defined as a half breadth, and then a time range of this doubledhalf breadth gives a timing window Tw.

Here, a coincidence counting processing in the coincidence countingprocessing part 6 will be described using FIG. 8 and FIG. 9 in detail.FIG. 8 shows a graph that illustrates a timing spectrum in the casewhere the difference of detection time based on the difference betweendecay times of scintillator arrays is not compensated. FIG. 9 shows agraph that illustrates a timing spectrum in the case where thedifference of detection time based on the difference between decay timesof scintillator arrays is compensated.

First, in conventional coincidence counting processing, the differenceof the detection time based on the difference of the decay times betweenthe scintillator array 18 a the and scintillator array 18 b is notcompensated as illustrated in FIG. 8. Four kinds of coincidence countingperformed in the coincidence counting processing part 6 have 4 kinds ofcombinations of: MD1t_(SF) and MD2t_(SF); MD1t_(SF) and MD2t_(SR);MD1t_(SR) and MD2t_(SF); and MD1t_(SR) and MD2t_(SR), for example, whenone γ ray detector 1 of two γ ray detectors 1 is defined as MD1 andanother γ ray detector 1 as MD2. Since MD1t_(SF) and MD2t_(SF), andMD1t_(SR) and MD2t_(SR) do not have difference in decay times ofscintillator arrays in the 4 kinds of combinations, no difference oftime will be given, leading to satisfactory coincidence counting.However, MD1t_(SR) and MD2t_(SF), and MD1t_(SR) and MD2t_(SR) havedifference of time based on the difference of decay times between thescintillator array 18 a and the scintillator array 18 b, and then a casemay occur where the count is not discriminated as an effectivecoincidence counting count, leading to possible drop of sensitivity.

Alternatively, in case of this embodiment, since the difference ofdetection time based on the difference in decay times between thescintillator array 18 a and the scintillator array 18 b is compensatedby the incidence timing compensation part 29, 4 kinds of coincidencecountings performed in the coincidence counting processing part 6 willbe within the timing window Tw as illustrated in FIG. 9, and therefore acounting loss of effective coincidence counting will not be generated,leading to avoidance of drop of sensitivity. Here, this timing window Twis stored in the timing window storage part 32 that is one function ofthe ROM9 of the FPGA7. The above-described coincidence countingprocessing part 6 is equivalent to the coincidence counting device. Theabove-described timing window storage part 32 is equivalent to thetiming window storing device.

Next, the operations will be described in order wherein a γ ray emittedfrom the subject M enters into the γ ray detector 1 and then issubjected to coincidence counting processing of the γ ray in thecoincidence counting processing part 6, with reference to FIG. 1, FIG.3, and FIG. 10. FIG. 10 illustrates a flow chart from an incidencetiming occurrence until incidence timing compensation processing iscarried out.

First, two γ rays emitted in a direction making approximately 180degrees with each other from a region of interest of the subject enterinto the opposing scintillator block 15 of the γ ray detectors 1, asillustrated in FIG. 1. The γ rays generate light in each of thescintillator 19 a of the scintillator array 18 a with a shorter decaytime for the emitted light pulse and the scintillator 19 b of thescintillator array 18 b with longer decay time for emitted light pulsethat constitute the scintillator block 15 as illustrated in FIG. 3. Thislight is guided to a light guide 16, and distributed to 4 PMTphotoelectric conversion plates of the photo multiplier tube (PMT) 17based on a position of incidence (X direction and Y direction of thescintillators 19 a and 19 b), reaching thereto. Furthermore, the lightsare converted into electrical signals (analog signals) in the photomultiplier tube 17, and are outputted to the amplification circuits 2 aand 2 b. The voltage of the analog signals are amplified in theamplification circuits 2 a and 2 b and are outputted to the incidencetiming calculation parts 4 a and 4 b and the A/D converters 3 a and 3 b.Furthermore, the analog signals inputted into the A/D converters 3 a and3 b are A/D converted into digital signals, and are temporarily storedin the A/D converted signal storage part 23.

Here, there will be described, with reference to FIG. 10, a flow fromcalculation of incidence timings t_(SF) and t_(SR) of the incidencetiming calculation parts 4 a and 4 b that have received analog signalsfrom the amplification circuits 2 a and 2 b to a processing in theincidence timing compensation part 29 of this incidence timing t_(SF)and t_(SR), out of operations from incidence of a γ ray emitted from thesubject M to the γ ray detector 1, up to the coincidence countingprocessing of the γ ray in the coincidence counting processing part 6.FIG. 10 is a flow chart illustrating the signal processing after theelectrical signals outputted from the photo multiplier tube 17 have beenoutputted from the amplification circuits 2 a and 2 b.

(Step S1)

The ARC20 of the incidence timing calculation parts 4 a and 4 bcalculate incidence timings t_(SF) and t_(SR) for analog signalsoutputted from the amplification circuits 2 a and 2 b based on inputs.Furthermore, the timing generator circuit 21 converts these incidencetimings t_(SF) and t_(SR) into digital signals, and outputs to theincidence timing storage part 22. Here, when the incidence timingst_(SF) and t_(SR) are stored in the incidence timing storage part 22(incidence timing t_(SF) and t_(SR) are generated), the flow proceeds toStep S2, and does not go to the next operation until the incidencetiming t_(SF) and t_(SR) are generated (operation of the step S1 isrepeated).

(Step S2)

The scintillator array identifying part 24 of the position arithmeticprocessing parts 5 a and 5 b reads the digital signals, after A/Dconversion, that are temporarily stored in the A/D converted signalstorage part 23 based on generation of the incidence timings t_(SF) andt_(SR), and the flow proceeds to Step S3.

(Step S3)

The scintillator array identifying part 24 of the position arithmeticprocessing parts 5 a and 5 b performs integration operation by addingsequentially the digital signals after A/D conversion. An intermediateadded value A_(T1) up to an intermediate T₁ (Fs×m: Fs representssampling interval of the A/D conversion, and m represents the number oftimes of addition), and a total added value A_(T2) up to the point oftime of emission end T₂ (Fs×n: n is the total number of times ofaddition) are acquired, and subsequently the flow proceeds to Step S4.

(Step S4)

Furthermore, the scintillator array identifying part 24 of the positionarithmetic processing parts 5 a and 5 b calculates an identified valuethat shows a value of (intermediate added value A_(T1))/(total addedvalue A_(T2)) from the intermediate added value A_(T1) and the totaladded value A_(T2), and it discriminates whether the identified value isa larger value or a smaller value as compared to an intermediate value Kstored in the intermediate value data table 27 based on emitted lightpulses emitted by each of the scintillators 19 a and 19 b of the two thescintillator arrays 18 a and 18 bs. Furthermore, it outputs signals thatshow these discriminated results to the incidence timing compensationpart 29 of the position arithmetic processing parts 5 a and 5 b. Here,when the identified value is smaller than the intermediate value K, theflow proceeds to Step S5, and conversely, when the identified value islarger than the intermediate value K, the flow proceeds to Step S6.

(Step S5)

When the identified value is smaller than the intermediate value K, thedecay time of the emitted light pulse of the scintillator arrayidentified by the scintillator array identifying part 24 is determinedto be longer. The incidence timing compensation part 29 of the positionarithmetic processing parts 5 a and 5 b discriminates that acompensation for substituting the incidence timing t_(SR) calculated bythe incidence timing calculation parts 4 a and 4 b by t_(SR)−Δt(incidence timing compensated value) is to be performed. Accordingly,the incidence timing compensation part 29 reads out the incidence timingt_(SR) stored in the incidence timing storage part 22, and performscompensation processing of substitution of this incidence timing t_(SR)to t_(SR)−Δt. Subsequently, it outputs a signal that illustrates thist_(SR)−Δt to the post-compensation incidence timing storage part 30, andthe signal is stored in the post-compensation incidence timing storagepart 30.

(Step S6)

When the identified value is larger than the intermediate value K, it isdetermined that the decay time of emitted light pulse of thescintillator array identified by the scintillator array identifying part24 is shorter. The incidence timing compensation part 29 of the positionarithmetic processing parts 5 a and 5 b discriminates that acompensation of the incidence timing t_(SR) calculated by the incidencetiming calculation parts 4 a and 4 b is not to be performed.Accordingly, the incidence timing compensation part 29 reads out theincidence timing t_(SR) stored in the incidence timing storage part 22,and outputs the signal that shows the incidence timing t_(SR) to thepost-compensation incidence timing storage part 30, as it is, withoutcompensation processing for this incidence timing t_(SR), to be storedin the post-compensation incidence timing storage part 30.

Next, the flow from the processing by the incidence timing compensationpart 29 of this incidence timings t_(SF) and t_(SR) to the coincidencecounting processing will be described with reference to FIG. 7, out ofthe operations from the incidence of the γ ray emitted from the subjectM into the γ ray detector 1 to the coincidence counting processing ofthe γ ray in the coincidence counting processing part 6. The coincidencecounting processing part 6 reads out the incidence timings t_(SF) andt_(SR) that have been stored in the post-compensation incidence timingstorage part every 128 ns. The coincidence counting processing part 6performs coincidence counting for the incidence timings t_(SF), t_(SR)from the 2γ ray detectors 1, in this case, 4 kinds of combinations. Whenthe difference of time calculated by 4 kinds of the combination existsin the timing window Tw (for example, 6 ns), the count concerned isdiscriminated as a valid coincidence count, and when not, it isdiscriminated as an invalid count.

According to the above-described nuclear medicine diagnosis equipment,the incidence timing compensation part 29 discriminates whether acompensation of an incidence timing calculated by the incidence timingcalculation parts 4 a and 4 b is to be performed or not corresponding tothe scintillator array identified by the scintillator array identifyingpart 24, and then compensation for the incidence timing is performedbased on the result of the discrimination. The difference of time ofdetections in the scintillator array 18 a and the scintillator array 18b generated based on the difference in decay times for the emitted lightpulse can be canceled by compensation, even in case of coincidencecounting using a scintillator array 18 a and a scintillator array 18 bhaving different decay time for an emitted light pulse. Accordingly, aprecise and accurate tomogram image providing a higher detectionsensitivity and avoidance of degradation of a reconstructed image may beobtained even in use, as a γ ray detector, of a scintillator havingdifferent decay time for the emitted light pulse.

Incidentally, identification of whether an identified value calculatedby the identified value calculating part 26 is a larger value or asmaller value may be performed based on sequential addition by theadding part 25 of the scintillator array identifying part 24. That is,the scintillator array identifying part 24 can identify whichscintillator array in the scintillator has emitted the light pulse.Furthermore, since the integral operation conventionally performed withan integrator may be replaced to an addition operation of sequentialaddition by the adder, reduction of parts marks and cost reduction maybe expected.

The present invention is not limited to the above-described embodiments,and modified implementation may be carried out as follows.

(1) Although a PET device has been adopted as an example and descriptionhas been given in the above-described embodiment, the present inventionis not limited to the PET devices, and may be applied to anyapparatuses, as long as it is a nuclear medicine equipment forperforming nuclear medicine diagnosis by coincidence counting of aradiation generated from a subject that has been given aradiopharmaceutical.

(2) In the above-described embodiments, application is also possible toa combination device of a nuclear medicine diagnosis equipment and anX-ray CT scanner like PET-CT provided with a PET device and an X-ray CTscanner.

(3) The timing window storage part 32 may store timing windows Twcorresponding to a combination of each of a plurality of scintillatorarrays. Accordingly, since different timing windows Tw are used withcombination of each of a plurality of scintillator arrays, coincidencecounting having high precision can be performed, and thereby influencesof occurrence coincidence counting, dispersion coincidence counting,etc. are eliminated, leading to higher definition image with lessnoises.

(4) The timing window storage part 32 may store the timing window Twcorresponding to respective combinations of the plurality ofscintillators. Accordingly, use of the different timing window Twaccording to combination of each of the plurality of scintillatorsallows coincidence counting having high precision, also allows reductionof influence of random coincidence counting, scattered coincidencecounting, etc., and also allows provision of higher definition imagehaving less noise.

(5) In the above-described embodiment, the incidence timing compensationpart 29 of the position arithmetic processing parts 5 a and 5 b performsan operation processing for substituting the incidence timing t_(SR)calculated by the incidence timing calculation parts 4 a and 4 b byt_(SR)−Δt (incidence timing compensated value), and the incidence timingt_(SR)−Δt and an incidence timing t_(SF) are to be stored in thepost-compensation incidence timing storage part 30 without compensationprocessing of the incidence timing t_(SF). However, the incidence timingcompensation part 29 of the position arithmetic processing parts 5 a and5 b may perform an operation processing for substituting the incidencetiming t_(SF) calculated by the incidence timing calculation parts 4 aand 4 b by t_(SF)+Δt, and the incidence timing t_(SF)+Δt and theincidence timing t_(SR) may be stored in the post-compensation incidencetiming storage part 30 without compensation processing of the incidencetiming t_(SR).

(6) In the above-described embodiment, a difference of time of the risetimes of the scintillator arrays 18 a and 18 b is stored in thecompensation data table 31 as an incidence timing value forcompensation. However, a difference of time of the different rise timesbetween the scintillator 19 s may be stored in the compensation datatable 31 as an incidence timing compensated value. Accordingly, thedifference of time of the different rise time between the scintillators19 can be compensated, leading to further improvement in detectionsensitivity.

(7) In the above-described embodiment, the difference of time of thedifferent rise times between the scintillator arrays 18 a and 18 b isbeforehand stored in the compensation data table 31 as an incidencetiming compensated value. However, an incidence timing compensated valuemay be calculated in real time, for example, using a simple linearfunction that uses an identified value as a variable utilizing anidentified value calculated by the identified value calculating device.

(8) In the above-described embodiment, the scintillator block 15 isconstituted of a scintillator array 18 a using Lu_(x)Y_(2-x)SiO₅ (LYSO)in an incidence side of a γ ray (front step) as a scintillator 19 ahaving a shorter decay time for an emitted light pulse and of ascintillator 19 b using Gd₂SiO₅ (GSO) with Ce concentration of 0.5 mol)having a longer decay time of the emitted light pulse in a light guide16 side (back step) as a scintillator 19 b having a longer decay timefor an emitted light pulse. Materials such as Gd₂SiO₅ (GSO) having Ceconcentration of 0.5 mol, Gd₂SiO₅ (GSO) having Ce concentration of 1.5mol, Lu₂SiO₅ (LSO), Lu_(x)Gd_(2-x)SiO₅ (LGSO), Lu_(x)Y_(2-x)SiO₅ (LYSO),Bi₄Ge₃O₁₂ (BGO), NaI, BaF₂, and CsF may be selected and may be used invarious combination for the scintillator 19 a of the scintillator array18 a and the scintillator 19 b of the scintillator array 18 b thatconstitute the scintillator block 15.

(9) In the above-described embodiment, although the scintillator block15 has been described as a block having combined two-layers (pieces) ofthe scintillator array 18 a and the scintillator array 18 b, they may becombination of two or more layers (pieces) in stead of the combinationof the two-layer (piece). Furthermore, although the number of thescintillators 19 a and 19 b of each scintillator has been described as acombination of 8 x 8, scintillators having a plurality of numbers ofcombination may also be used.

(10) In the above-described embodiment, although the photo multipliertube 17 was exemplified as a photodetector, photodetectors other thanthe embodiment, for example, photo-diodes, avalanche photo-diodes, etc.,may be used.

1. A nuclear medicine diagnosis equipment, comprising: a scintillatorblock having a plurality of two-dimensionally and closely arrangedscintillators, the scintillator block having a plurality of opticallycombined scintillator arrays in a depth direction of an incident γ raywith different decay times for an emitted light pulse; a photodetectorfor converting an emitted light pulse emitted in the scintillator blockinto an electrical signal; an incidence timing calculating device forcalculating an incident timing in the scintillator array for theelectrical signal outputted from the photodetector; a scintillator arrayidentifying device for identifying a scintillator array, in a pluralityof arrays, that has received the electrical signal outputted from thephotodetector; and an incidence timing compensation device fordiscriminating whether compensation for an incidence timing calculatedby the incidence timing calculating device is to be done or notcorresponding to a scintillator array identified by the scintillatorarray identification device, and for compensating the incidence timingbased on a result of discrimination.
 2. The nuclear medicine diagnosisequipment according to claim 1, comprising an A/D converter forconverting an analog signal in a form of an electrical signal outputtedfrom a photodetector into a digital signal, the scintillator arrayidentifying device comprising: an adding device for sequentially addingdigital signals converted by the A/D converter; an identified valuecalculating device for calculating an identified value giving a valueobtained by division of an intermediate added value by a total addedvalue by using an intermediate added value obtained, in the addingdevice, by addition of the digital signals from a point of time ofemission start of the emitted light pulse that has been emitted in thescintillator block to a certain middle point of time in the course of apoint of time of emission end, and a total added value obtained, in theadding device, by addition of a digital signal from a point of time ofemission start to a point of time of emission end of the emitted lightpulse that has been emitted in the scintillator block in the addingdevice; and a discriminating device for discriminating whether theidentified value calculated by the identified value calculating deviceis a larger value or a smaller value as compared with an intermediatevalue between the identified values of each scintillator arraycalculated by the identified value calculating device.
 3. The nuclearmedicine diagnosis equipment according to claim 1, comprising: acoincidence counting device for performing coincidence counting using anincidence timing compensated by the incidence timing compensation deviceand an incidence timing discriminated as not compensated by theincidence timing compensation device; and a timing window storing devicefor storing a timing window showing a predetermined range for performingcoincidence counting by the coincidence counting device as a timingwindow corresponding to a combination of each of the plurality ofscintillator arrays.
 4. The nuclear medicine diagnosis equipmentaccording to claim 1, comprising: a coincidence counting device forperforming coincidence counting using an incidence timing compensated bythe incidence timing compensation device and an incidence timingdiscriminated as not compensated by the incidence timing compensationdevice; and a timing window storing device for storing a timing windowshowing a predetermined range for performing coincidence counting by thecoincidence counting device as a timing window corresponding to acombination of each of the plurality of scintillators.
 5. The nuclearmedicine diagnosis equipment according to claim 1, comprising a lightguide for optically coupling the scintillator block and thephotodetector.
 6. The nuclear medicine diagnosis equipment according toclaim 1, wherein the a plurality of scintillator arrays is constitutedby either one of scintillators of Gd₂SiO₅ (GSO) having Ce concentrationof 0.5 mol, Gd₂SiO₅ (GSO) having Ce concentration of 1.5 mol, Lu₂SiO₅(LSO), Lu_(x)Gd_(2-x)SiO₅ (LGSO), Lu_(x)Y_(2-x)SiO₅ (LYSO), Bi₄Ge₃O₁₂(BGO), NaI, BaF₂, and CsF.
 7. The nuclear medicine diagnosis equipmentaccording to claim 1, wherein the photodetector is made of a photomultiplier tube.
 8. The nuclear medicine diagnosis equipment accordingto claim 1, wherein the photodetector is made of a photo-diode.
 9. Thenuclear medicine diagnosis equipment according to claim 1, wherein thephotodetector is made of an avalanche photodiode.